The present invention generally relates to radio frequency (RF) receiver coils for magnetic resonance applications.
Magnetic resonance imaging (MRI) detects the faint nuclear magnetic resonance (NMR) signals given off by protons or other nuclei in the presence of a strong magnetic field after excitation with a RF signal. The NMR signals are detected using loop antennas termed “coils” or resonators.
Coils for magnetic resonance applications are typically operated at the Larmor frequency. The Larmor frequency depends on the strength of the basic magnetic field of the magnetic resonance system and on the chemical element whose excited spin is being detected. For hydrogen (which is the most frequent case), the gyromagnetic ratio is approximately 42.57 MHz/Tesla (1 Tesla=10000 gaus). During operation of a coil at resonance, a current oscillates with a resonance frequency in the conductor element. This current is particularly high when the conductor element is tuned to the resonance frequency. Thus, in the ideal case, the Larmor frequency corresponds to the resonance frequency of the resonator or coil.
NMR signals are, relatively speaking, extremely faint and therefore “local coils” or “surface coils” may be designed to be placed in close proximity to the region of interest of the imaged object. The size of the local coils is kept small to allow them to be easily fit to the patient on the MRI patient table. Importantly, the small area of loops of the local coil provides improved signal strength relative to received noise. The local coils are in contrast to the whole body coil typically present in an MRI machine and useful for obtaining broad survey scans of the patient.
The small size of a local coil generally limits the volume over which the coil is sensitive. For imaging large areas of the body, for example, neurovascular imaging of the head, neck, and lower spine, the whole body coil with its lower signal to noise ratio (SNR) must be used. Alternatively, coverage of this region can be obtained by using several local coils, taking multiple images of the patient and changing or repositioning the local coil in between images. This latter approach is time consuming and impractical in many situations. Additionally, these approaches may still suffer from lower SNR than is desired for imaging.
Mutli-layer conductors fashioned as resonators or coils for magnetic resonance imaging are also known, and examples include U.S. Pat. Nos. 7,579,835 and 7,579,836. However, as with single element coils, the use of multi-layer conductors in an MR coil does not, in itself, lead to improvements in SNR and uniformity over a range of clinically relevant loading conditions. In addition, when attempting to maintain SNR between subjects that present different loading conditions, a single coil would need to be re-tuned for each subject, or a number of single coils that were each tuned to different loading conditions would be needed. Alternatively, a broadband coil could be used to provide greater uniformity across loading conditions, but these broadband coils suffer from substantially lower SNR than could be achieved with single coils.